Fragmented polymeric compositions and methods for their use

ABSTRACT

Cross-linked hydrogels comprise a variety of biologic and non-biologic polymers, such as proteins, polysaccharides, and synthetic polymers. Such hydrogels preferably have no free aqueous phase and may be applied to target sites in a patient&#39;s body by extruding the hydrogel through an orifice at the target site. Alternatively, the hydrogels may be mechanically disrupted and used in implantable articles, such as breast implants. When used in vivo, the compositions are useful for controlled release drug delivery, for inhibiting post-surgical spinal and other tissue adhesions, for filling tissue divots, tissue tracts, body cavities, surgical defects, and the like.

The present application is a continuation-in-part of Application Ser.No. 08/903,674, filed on Jul. 31, 1997, which was a continuation-in-partof provisional Application No. 60/050,437, filed on Jun. 18, 1997, andwas a continuation-in-part of Application Ser. No. 08/704,852, filed onAug. 27, 1996, abandoned. The full disclosures of each of theseapplications are incorporated herein by reference.

BACKGROUND OF THE INVENTION Field of the Invention

The present invention relates generally to biocompatible cross-linkedpolymeric compositions and to the use of such compositions for thecontrolled delivery of aqueous agents to target sites.

It has long been recognized that tablets, capsules, and injections arenot the optimum route of drug delivery for all purposes. Theseconventional routes often require frequent and repeated doses, resultingin a "peak and valley" pattern of therapeutic agent concentration. Sinceeach therapeutic agent has a therapeutic range above which it is toxicand below which it is ineffective, a fluctuating therapeutic agentconcentration may cause alternating periods of ineffectiveness andtoxicity. For this reason, a variety of "controlled release" drugformulations and devices have been proposed for maintaining thetherapeutic agent level within the desired therapeutic range for theduration of treatment. Using a polymeric carrier is one effective meansto deliver the therapeutic agent locally and in a controlled fashion. Inaddition to controlled levels, such systems often require less totaldrug and minimize systemic side effects.

Polymeric carriers may be biodegradable or non-biodegradable. For anon-biodegradable matrix, the steps leading to release of thetherapeutic agent are water diffusion into the matrix, dissolution ofthe therapeutic agent, and out-diffusion of the therapeutic agentthrough the channels of the matrix. As a consequence, the mean residencetime of the therapeutic agent existing in the soluble state is longerfor a non-biodegradable matrix than for a biodegradable matrix where along passage through the channels is no longer required. Since manypharmaceuticals have short half-lives, there is a significant chancethat the therapeutic agent may be decomposed or inactivated inside thenon-biodegradable matrix before it can be released. The risk isparticularly significant for many biological macromolecules and smallerpolypeptides, since these molecules are generally unstable in buffer andhave low permeability through polymers. In fact, in a non-biodegradablematrix, many bio-macromolecules will aggregate and precipitate, cloggingthe channels necessary for diffusion out of the carrier matrix.

These concerns are largely alleviated by using a biodegradablecontrolled release matrix. Biodegradable polymers release containeddrugs as the matrix is consumed or biodegraded during therapy. Thepolymer is usually selected to breakdown into subunits which arebiocompatible with the surrounding tissue. The persistence of abiodegradable polymer in vivo depends on its molecular weight and degreeof cross-linking, the higher the molecular weights and degrees ofcross-linking resulting in a longer life. Common biodegradable polymersinclude polylactic acid (PLA, also referred to as polylactide),polyglycolic acid (PGA), copolymers of PLA and PGA, polyamides, andcopolymers of polyamides and polyesters. PLA undergoes hydrolyticde-esterification to lactic acid, a normal product of muscle metabolism.PGA is chemically related to PLA and is commonly used for absorbablesurgical sutures, as in the PLA/PGA copolymer. However, the use of PGAin controlled-release implants has been limited due to its lowsolubility in common solvents and subsequent difficulty in fabricationof devices.

An additional advantage of biodegradable drug delivery carriers is theelimination of the need for surgical removal after it has fulfilled itsmission. Additional advantages include: 1) the ability to controlrelease rate through variation of the matrix composition; 2) the abilityto implant at sites difficult or impossible for retrieval; 3) animproved ability to deliver unstable therapeutic agents. This last pointis of particular importance in light of the advances in molecularbiology and genetic engineering which have lead to the commercialavailability of many potent biological macromolecules. Suchmacromolecules usually have short in vivo half-lives and low GI tractabsorption which often render them unsuitable for conventional oral orintravenous administration.

Ideally, a biodegradable therapeutic agent delivery system would simplyconsist of a solution, suspension, or dispersion of the drug in apolymer matrix. The therapeutic agent is released as the polymericmatrix decomposes, or biodegrades into soluble products which areexcreted from the body. Unfortunately, the ability to design idealbiodegradable delivery systems is limited by many characteristics of thepolymers, including weak mechanical strength, unfavorable degradationcharacteristics, toxicity, inflexibility, fabrication difficulty, andthe like. Although known biodegradable polymers have a broad range ofpotential utility, there is no one single material available that couldsatisfy all requirements imposed by different applications. Accordingly,there continues to be need to develop new biodegradable polymers.

U.S. Patent Nos. 5,672,336 and 5,196,185 describe a wound dressingcomprising a micro-particulate fibrillar collagen having a particle sizeof 0.5-2.0 μm. This composition generally comprises an aqueous phase anddoes not form a hydrogel as described in the present invention. U.S.Pat. No. 5,698,213 describes a cross-linked aliphatic poly-esterhydrogel useful as an absorbable surgical device and drug deliveryvehicle. U.S. Pat. No. 5,674,275 describes an acrylate or methacrylatebased hydrogel adhesive. U.S. Pat. No. 5,306,501 describes apolyoxyalkylene based thermoreversible hydrogel useful as a drugdelivery vehicle.

U.S. Pat. No. 4,925,677 and U.S. Pat. No. 5,041,292 describe a hydrogelcomprising a protein component cross-linked with a polysaccharide ormucopolysaccharide and useful as a drug delivery vehicle.

For these reasons, it would be desirable to provide improvedcompositions, methods, and kits for delivering biological macromoleculeand other drugs to target body sites. In particular, it would bedesirable to provide compositions which are compatible with a widevariety of drugs either in solution or in suspension, particularly drugspresent in an aqueous carrier. Still more preferably, the compositionsshould be in the form of hydrogels which are biocompatible and whichpermit substantial control or "programming" of the releasecharacteristics, including release rate, composition persistence, drugcarrying capacity, product delivery characteristics (such asinjectability), and the like. In addition to drug delivery and release,the products, methods, and kits of the present invention should beadaptable for localizing active agents at a target site, where theactive agents can provide biological activity even prior to release fromthe product matrix. At least some of these objectives will be met by theembodiments of the invention described hereinafter.

Biodegradable injectable drug delivery polymers are described in U.S.Pat. No. 5,384,333 and by Jeong et al. (1997) "Nature," 388:860-862.Biodegradable hydrogels for controlled released drug delivery aredescribed in U.S. Pat. No. 4,925,677. Resorbable collagen-based drugdelivery systems are described in U.S. Pat. Nos. 4,347,234 and4,291,013. Aminopolysaccharide-based biocompatible films for drugdelivery are described in U.S. Pat. Nos. 5,300,494 and 4,946,870. Watersoluble carriers for the delivery of taxol are described in U.S. Pat.No. 5,648,506.

Polymers have been used as carriers of therapeutic agents to effect alocalized and sustained release (Langer, et al., Rev. Macro. Chem.Phys., C23 (1), 61, 1983; Controlled Drug Delivery, Vol. I and II,Bruck, S.D., (ed.), CRC Press, Boca Raton, Fla., 1983; Leong et al.,Adv. Drug Delivery Review, 1:199, 1987). These therapeutic agentdelivery systems simulate infusion and offer the potential of enhancedtherapeutic efficacy and reduced systemic toxicity.

Other classes of synthetic polymers which have been proposed forcontrolled release drug delivery include polyesters (Pitt, et al., inControlled Release of Bioactive Materials, R. Baker, Ed., AcademicPress, New York, 1980); polyamides (Sidman, et al., Journal of MembraneScience, 7:227, 1979); polyurethanes (Maser, et al., Journal of PolymerScience, Polymer Symposium, 66:259, 1979); polyorthoesters (Heller, etal., Polymer Engineering Scient, 21:727, 1981); and polyanhydrides(Leong, et al., Biomaterials, 7:364, 1986).

Collagen-containing compositions which have been mechanically disruptedto alter their physical properties are described in U.S. Pat. Nos.5,428,024; 5,352,715; and 5,204,382. These patents generally relate tofibrillar and insoluble collagens. An injectable collagen composition isdescribed in U.S. Pat. No. 4,803,075. An injectable bone/cartilagecomposition is described in U.S. Pat. No. 5,516,532. A collagen-baseddelivery matrix comprising dry particles in the size range from 5 μm to850 μm which may be suspended in water and which has a particularsurface charge density is described in WO 96/39159. A collagenpreparation having a particle size from 1 μm to 50 μm useful as anaerosol spray to form a wound dressing is described in U.S. Pat. No.5,196,185. Other patents describing collagen compositions include U.S.Pat. Nos. 5,672,336 and 5,356,614.

A polymeric, non-erodible hydrogel that may be cross-linked and injectedvia a syringe is described in WO 96/06883.

The following pending applications, assigned to the assignee of thepresent application, contain related subject matter: U.S. Ser. No.08/903,674, filed on Jul. 31, 1997; U.S. Ser. No. 60/050,437, filed onJun. 18, 1997; U.S. Ser. No. 08/704,852, filed on Aug. 27, 1996; U.S.Ser. No. 08/673,710, filed Jun. 19, 1996; U.S. Ser. No. 60/011,898,filed Feb. 20, 1996; U.S. Ser. No. 60/006,321, filed on Nov. 7, 1996;U.S. Ser. No. 60/006,322, filed on Nov. 7, 1996; U.S. Ser. No.60/006,324, filed on Nov. 7, 1996; and U.S. Ser. No. 08/481,712, filedon Jun. 7, 1995. The full disclosures of each of these applications isincorporated herein by reference.

SUMMARY OF THE INVENTION

The present invention provides improved biocompatible polymericcompositions and methods for applying such compositions at target sitesin a patient's body. The methods and compositions will be particularlyuseful for delivering drugs and other active agents, such as biologicalmacromolecules, polypeptides, oligopeptides, nucleic acids, smallmolecule drugs, and the like. The compositions will comprisebiocompatible, cross-linked hydrogels, as described in more detailbelow, and the drug or other biologically active agent will typically beincorporated into the composition as an aqueous solution, suspension,dispersion, or the like. The drugs may be incorporated into thecompositions prior to packaging, immediately prior to use, or even asthe compositions are being applied to the target site. Afterintroduction to the target site, the drug will usually be released overtime as the composition degrades. In some instances, however, the drugor other biological agent may display activity while still incorporatedor entrapped within the hydrogel. For example, the compositions andmethods may find specific use in stopping or inhibiting bleeding(hemostasis), particularly when combined with a suitable hemostaticagent, such as thrombin, fibrinogen, clotting factors, and the like.

The compositions will have other uses as well, such as tissuesupplementation, particularly for filling soft and hard tissue regions,including divots, tracts, body cavities, etc., present in muscle, skin,epithelial tissue, connective or supporting tissue, nerve tissue,ophthalmic and other sense organ tissue, vascular and cardiac tissue,gastrointestinal organs and tissue, pleura and other pulmonary tissue,kidney, endocrine glands, male and female reproductive organs, adiposetissue, liver, pancreas, lymph, cartilage, bone, oral tissue, andmucosal tissue. The compositions of the present invention will be stillfurther useful for filling soft implantable devices, such as breastimplants, where the material will be protected from cellular or enzymedegradation by a protective barrier or cover. The compositions willadditionally be useful in other procedures where it is desirable to filla confined space with a biocompatible and resorbable polymeric material.Additionally, the compositions may also find use for inhibiting theformation of tissue adhesions, such as spinal tissue adhesions,following surgery and traumatic injury.

The compositions of the present invention comprise a biocompatible,molecular cross-linked hydrogel. Usually the compositions will havesubstantially no free aqueous phase as defined herein below. Thehydrogel is resorbable and fragmented, i.e. comprises small subunitshaving a size and other physical properties which enhance theflowability of the hydrogel (e.g. the ability to be extruded through asyringe) and the ability of the hydrogel to otherwise be applied ontoand conform to sites on or in tissue, including tissue surfaces anddefined cavities, e.g. intravertebral spaces, tissue divots, holes,pockets, and the like. In particular, the fragmented subunits are sizedto permit them to flow when the compositions are subjected to stressesabove a threshold level, for example when extruded through an orifice orcannula, when packed into a delivery site using a spatula, when sprayedonto the delivery site, or the like. The threshold stresses aretypically in the range from 3×10⁴ Pa to 5×10⁵ Pa. The compositions,however, will remain generally immobile when subjected to stresses belowthe threshold level.

By "biocompatible," it is meant that the compositions will be suitablefor delivery to and implantation within a human patient. In particular,the compositions will be non-toxic, non-inflammatory (or will display alimited inflammatory effect which is not inconsistent with theirimplantation), and be free from other adverse biological effects.

By "biodegradable," it is meant that the compositions will degrade andbreakdown into smaller molecular subunits that will be resorbed and/oreliminated by the body over time, preferably within the time limits setforth below.

By "substantially free of an aqueous phase" it is meant that thecompositions will be fully or partially hydrated, but will not behydrated above their capacity to absorb water. In particular, a test fordetermining whether a composition has a free aqueous phase is set forthin Example 8 below. Hydrogels which are substantially free of an aqueousphase should release less than 10% by weight aqueous phase whensubjected to a 10 lb. force in the test, preferably releasing less than5% by weight, and more preferably less than 1% by weight, and morepreferably releasing no discernable aqueous phase and displaying nocollapse.

The compositions may be dry, partially hydrated or fully hydrateddepending on the extent of hydration. The fully hydrated material willhold from about 400% to about 5000% water or aqueous buffer by weight,corresponding to a nominal increase in diameter or width of anindividual particle of subunit in the range from approximately 50% toapproximately 500%, usually from approximately 50% to approximately250%. Thus, the size of particles in the dry powder starting material(prior to hydration) will determine the partially or fully hydrated sizeof the subunit (depending on the factors described below). Exemplary andpreferred size ranges for the dry particles and fully hydrated subunitsare as follows:

    ______________________________________                                        Particle/Subunit Size                                                                    Exemplary Range                                                                         Preferred Range                                          ______________________________________                                        Dry Particle 0.01 mm-1.5 mm                                                                            0.05 mm-1 mm                                         Fully Hydrated                                                                             0.02 mm-3 mm                                                                              0.1 mm-1.5 mm                                        Hydrogel Subunit                                                              ______________________________________                                    

Compositions of the present invention will usually be in the form of adry powder, a partially hydrated hydrogel, or a fully hydrated hydrogel.The dry powder (having a moisture content below 20% by weight) will beuseful as a starting material for preparation of the hydrogels, asdescribed below. The partially hydrated hydrogels are useful forapplications where it is desired that the material further swell uponapplication to a moist target site, e.g. a tissue divot. The fullyhydrated forms will be useful for applications where in situ swelling isnot desired, such as in the spinal column and other areas where nervesand other sensitive structures are present.

The dimensions of the subunits may be achieved in a variety of ways. Forexample, a cross-linked hydrogel having dimensions larger than thetarget range (as defined below) may be mechanically disrupted at avariety of points during the production process. In particular, thecomposition may be disrupted (1) before or after cross-linking of apolymer starting material and (2) before or after hydration of thecross-linked or non-cross-linked polymer starting material, e.g. as afully or partially hydrated material or as a dry particulate powder. Theterm "dry" will mean that the moisture content is sufficiently low,typically below 20% by weight water, so that the powder will befree-flowing and that the individual particles will not aggregate. Theterm "hydrated" will mean that the moisture content is sufficientlyhigh, typically above 50% of the equilibrium hydration level, usually inthe range from 70% to 95% of the equilibrium hydration level, so thatthe material will act as a hydrogel.

Mechanical disruption of the polymer material in the dry state ispreferred in cases where it is desired to control the particle sizeand/or particle size distribution. It is easier to control comminutionof the dry particles than the hydrated hydrogel materials, and the sizeof the resulting reduced particles is thus easier to adjust. Conversely,mechanical disruption of the hydrated, cross-linked hydrogels isgenerally simpler and involves fewer steps than does comminution of adry polymer starting material. Thus, the disruption of hydratedhydrogels may be preferred when the ultimate hydrogel subunit sizeand/or size distribution is less critical.

In a first exemplary production process, a dry, non-cross-linked polymerstarting material, e.g. dry gelatin powder, is mechanically disrupted bya conventional unit operation, such as homogenization, grinding,coacervation, milling, jet milling, and the like. The powder will bedisrupted sufficiently to achieve dry particle sizes which producehydrogel subunit sizes in the desired ranges when the product ispartially or fully hydrated. The relationship between the dry particlesize and the fully hydrated subunit size will depend on the swellabilityof the polymeric material, as defined further below.

Alternatively, a particulate polymeric starting material may be formedby spray drying. Spray drying processes rely on flowing a solutionthrough a small orifice, such as a nozzle, to form droplets which arereleased into a counter-current or co-current gas stream, typically aheated gas stream. The gas evaporates solvent from the liquid startingmaterial, which may be a solution, dispersion, or the like. Use of spraydrying to form a dry powder starting material is an alternative tomechanical disruption of the starting material. The spray dryingoperation will usually produce a non-cross-linked dry powder productwhich is spherical in shape with a generally uniform particle size. Theparticles may then be cross-linked, as described below.

In many instances, the mechanical disruption operation can be controlledsufficiently to obtain both the particle size and particle sizedistribution within a desired range. In other cases, however, where moreprecise particle size distributions are required, the disrupted materialcan be further treated or selected to provide the desired particle sizedistribution, e.g. by sieving, aggregation, or the like. Themechanically disrupted polymeric starting material is then cross-linkedas described in more detail below, and dried.

The dried material may be the desired final product, where it may berehydrated and swollen immediately prior to use. Alternatively, themechanically disrupted, cross-linked material may be rehydrated, and therehydrated material packaged for storage and subsequent use. Particularmethods for packaging and using these materials are described below.

Where the subunit size of the fragmented hydrogel is less important, thedried polymeric starting material may be hydrated, dissolved, orsuspended in a suitable buffer and cross-linked prior to mechanicaldisruption. Mechanical disruption of the pre-formed hydrogel willtypically be achieved by passing the hydrogel through an orifice, wherethe size of the orifice and force of extrusion together determine theparticle size and particle size distribution. While this method is oftenoperationally simpler than the mechanical disruption of dry polymericparticles prior to hydration and cross-linking, the ability to controlthe hydrogel particle size is much less precise.

In a particular aspect of the mechanical disruption of pre-formedhydrogels, the hydrogels may be packed in a syringe or other applicatorprior to mechanical disruption. The materials will then be mechanicallydisrupted as they are applied through the syringe to the tissue targetsite, as discussed in more detail below. Alternatively, a non-disrupted,cross-linked polymeric material may be stored in a dry form prior touse. The dry material may then be loaded into a syringe or othersuitable applicator, hydrated within the applicator, and mechanicallydisrupted as the material is delivered to the target site, againtypically being through an orifice or small tubular lumen.

The polymer will be capable of being cross-linked and of being hydratedto form a hydrogel, as described in more detail below. Exemplarypolymers include proteins selected from gelatin, collagen (e.g. solublecollagen), albumin, hemoglobin, fibrinogen, fibrin, fibronectin,elastin, keratin, laminin, casein and derivatives and combinationsthereof. Alternatively, the polymer may comprise a polysaccharide, suchas a glycosaminoglycan (e.g., hyaluronic acid or chondroitin sulfate), astarch derivative, a cellulose derivative, a hemicellulose derivative,xylan, agarose, alginate, chitosan, and combinations thereof. As afurther alternative, the polymer may comprise a non-biologichydrogel-forming polymer, such as polyacrylates, polymethacrylates,polyacrylamides, polyvinyl polymers, polylactide-glycolides,polycaprolactones, polyoxyethylenes, and derivatives and combinationsthereof.

Cross-linking of the polymer may be achieved in any conventional manner.For example, in the case of proteins, cross-linking may be achievedusing a suitable cross-linking agent, such as an aldehyde, sodiumperiodate, epoxy compounds, and the like. Alternatively, cross-linkingmay be induced by exposure to radiation, such as γ-radiation or electronbeam radiation. Polysaccharides and non-biologic polymers may also becross-linked using suitable cross-linking agents and radiation.Additionally, non-biologic polymers may be synthesized as cross-linkedpolymers and copolymers. For example, reactions between mono- andpoly-unsaturated monomers can result in synthetic polymers havingcontrolled degrees of cross-linking. Typically, the polymer moleculeswill each have a molecular weight in the range from 20 kD to 200 kD, andwill have at least one link to another polymer molecule in the network,often having from 1 to 5 links, where the actual level of cross-linkingis selected in part to provide a desired rate of biodegradability in theranges set forth below.

The extent of cross-linking of the polymer has an effect on severalfunctional properties of the hydrogel including extrudability,adsorptiveness of surrounding biological fluids, cohesiveness, abilityto fill space, swelling ability and ability to adhere to the tissuesite. The extent of cross-linking of the polymeric hydrogel compositionmay be controlled by adjusting the concentration of cross-linking agent,controlling exposure to cross-linking radiation, changing the relativeamounts of mono- and poly-unsaturated monomers, varying reactionconditions, and the like. Typically, the degree of cross-linking iscontrolled by adjusting the concentration of cross-linking agent.

Exposure to radiation, such as γ-radiation, may also be carried out inorder to sterilize the compositions before or after packaging. When thecompositions are composed of radiation-sensitive materials, it will benecessary to protect the compositions from the undesirable effects ofsterilizing radiation. For example, in some cases, it will be desirableto add a stabilizer, such as ascorbic acid, in order to inhibitdegradation and/or further excessive cross-linking of the materials byfree radical mechanisms.

The hydrogel compositions of the present invention will typically have asolids content in the range from 1% by weight to 70% by weight.Optionally, the compositions may comprise at least one plasticizer asdescribed in more detail below. Suitable plasticizers includepolyethylene glycols, sorbitol, glycerol, and the like.

The equilibrium swell of the cross-linked polymers of the presentinvention may range from 400% to 5000%, 400% to 3000%, 400% to 2000%,usually ranging from 400% to 1300%, preferably being from 500% to 1100%,depending on its intended use. Such equilibrium swell may be controlledby varying the degree of cross-linking, which in turn is achieved byvarying the cross-linking conditions, such as the type of cross-linkingmethod, duration of exposure of a cross-linking agent, concentration ofa cross-linking agent, cross-linking temperature, and the like.

Gelatin-containing hydrogels having equilibrium swell values from about400% to 1300% were prepared and are described in the Experimentalsection hereinafter. It was found that materials having differingequilibrium swell values perform differently in different applications.For example, the ability to inhibit bleeding in a liver divot model wasmost readily achieved with cross-linked gelatin materials having a swellin the range from 600% to 950%, often from 700% to 950%. For a femoralartery plug, lower equilibrium swell values in the range from 500% to600% were more successful. Thus, the ability to control cross-linkingand equilibrium swell allows the compositions of the present inventionto be optimized for a variety of uses.

In addition to equilibrium swell, it is also important to control thehydration of the material immediately prior to delivery to a targetsite. Hydration is defined as the percentage of water contained by thehydrogel compared to that contained by the hydrogel when its fullysaturated, that is, at its equilibrium swell. A material with 0%hydration will be non-swollen. A material with 100% hydration will be atits equilibrium water content and fully swollen. Hydrations between 0%and 100% will correspond to swelling between the minimum and maximumamounts. As a practical matter, many dry, non-swollen materialsaccording to the present invention will have some residual moisturecontent, usually below 20% by weight, more usually from 8% to 15% byweight. When the term "dry" is used herein, it will specify materialshaving a low moisture content, usually below 20%, often below 10%, andfrequently below 5% by weight, where the individual particles are freeflowing and generally non-swollen.

Hydration can be adjusted very simply by controlling the amount ofaqueous buffer added to a dry or partially hydrated cross-linkedmaterial prior to use. Usually, at a minimum, it will be desirable tointroduce sufficient aqueous buffer to permit extrusion through asyringe or other delivery device. In other cases, however, it may bedesirable to utilize a spatula or other applicator for delivering lessfluid materials. The intended use will also help determine the desireddegree of hydration. In cases where it is desired to fill or seal amoist cavity, it is generally desirable to employ a partially hydratedhydrogel which can swell and fill the cavity by absorbing moisture fromthe target site. Conversely, fully or substantially fully hydratedhydrogels are preferred for application in the brain, near the spine,and to target sites near nerves and other sensitive body structureswhich could be damaged by post-placement swelling. It would also bepossible to prepare the hydrogel compositions of the present inventionwith excess buffer, resulting in a two-phase composition having a fullyhydrated hydrogel and a free buffer phase.

A preferred hydrogel material according to the present invention is agelatin which has been cross-linked to achieve from 600% to 950%,usually 700% to 950% swell at equilibrium hydration. The material willbe disrupted to have a hydrogel particle size in the range from 0.01 mmto 1.75 mm, preferably 0.05 mm to 1.0 mm, often 0.05 mm to 0.75 mm, andfrequently between 0.05 mm and 0.5 mm, and will preferably be hydratedat a level sufficient to achieve 70% to 100% of the equilibrium swellprior to application to the site.

In some cases, the hydrogel compositions of the present invention maycontain a combination of two or more different materials, e.gcombinations of proteins and polysaccharides and/or non-biologicpolymers, as well as combinations of two or more individual materialsfrom each of the polymer types, e.g. two or more proteins,polysaccharides, etc.

The polymeric compositions of the present invention may also comprisecombinations of the disrupted, cross-linked polymer hydrogels describedabove and non-cross-linked polymeric materials. The disrupted,cross-linked polymeric hydrogels consist of a plurality of subunitshaving a size determined by preparation method. The size is selected tobe useful for packing a confined volume, having both the flowability andthe rate of biodegradability described in the Experimental sectionbelow. The discrete nature of the cross-linked subunits, however, willleave void areas which may be filled by combination with anon-cross-linked polymeric material. The non-cross-linked polymeric orother filler material may comprise any of the polymeric materials listedabove, and may optionally but not necessarily be the same polymericmaterial which has been cross-linked to form the cross-linkedmechanically disrupted hydrogel. The relative amounts of cross-linkedpolymer and non-cross-linked polymer may vary, typically having a weightratio in the range from 20:1 to 1:1 (cross-linkedpolymer:non-cross-linked polymer), usually in the range from 10:1 to2:1, preferably from 5:1 to 2:1.

The hydrogels of the present application may be applied using a syringe,a spatula, a brush, a spray, manually by pressure, or by any otherconventional technique. Usually, the hydrogels will be applied using asyringe or similar applicator capable of extruding the hydrogel throughan orifice, aperture, needle, tube, or other passage to form a bead,layer, or similar portion of material. Mechanical disruption of thehydrogels can occur as the hydrogel is extruded through an orifice inthe syringe or other applicator, typically having a size in the rangefrom 0.01 mm to 5.0 mm, preferably 0.5 mm to 2.5 mm. Preferably,however, the polymeric hydrogel will be initially prepared from a powderhaving a desired particle size (which upon hydration yields hydrogelsubunits of the requisite size) or will be partially or entirelymechanically disrupted to the requisite size prior to a final extrusionor other application step.

The compositions may be applied at varying degrees of hydration, usuallybut not necessarily being at least partially hydrated. If applied in anon-hydrated form, the compositions will swell to their full equilibriumswell value, i.e. from about 400% to about 5000% as set forth above.When applied at their full hydration, the compositions will displaysubstantially equilibrium hydration and little or no swelling whenapplied to tissue. Swelling of the non-hydrated and partially hydratedcompositions results from absorption of moisture from the tissue andsurroundings to which the composition is being applied.

The present invention still further provides kits comprising any of thehydrated or non-hydrated hydrogel materials described above incombination with written instructions for use (IFU) which set forth anyof the methods described above for applying the hydrogel onto a targetsite on tissue. The composition and written instructions will beincluded together in a conventional container, such as a box, jar,pouch, tray, or the like. The written instructions may be printed on aseparate sheet of paper or other material and packaged on or within thecontainer or may be printed on the container itself. Usually, thecomposition(s) will be provided in a separate sterile bottle, jar, vial,or the like. When the hydrogel material is non-hydrated, the kit mayoptionally include a separate container with a suitable aqueous bufferfor hydration. Other system components such as the applicator, e.g.syringe, may also be provided.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 illustrates the application of the molecular cross-linkedpolymeric hydrogel of the present invention to a surgically createddefect in the vertebral body for preventing post-surgical spinaladhesions.

FIGS. 2A and 2B illustrate application of the molecular cross-linkedpolymeric hydrogel compositions of the present invention to a defect insoft tissue, where the treated region is optionally covered with aprotective patch after the defect is filled with the polymericcomposition.

FIGS. 3A and 3B illustrate use of the molecular cross-linked polymericcompositions of the present invention for filling a percutaneous tissuepenetration to a blood vessel, such as a tissue tract formed as part ofan intravascular catheterization procedure.

FIG. 4 illustrates a kit comprising a sterile package for an applicatorcontaining the molecular cross-linked polymeric composition of thepresent invention.

FIG. 5 illustrates the correlation between percent swell and the percentsolids of the polymeric hydrogel.

DESCRIPTION OF THE SPECIFIC EMBODIMENTS

Compositions according to the present invention comprise resorbablebiocompatible molecular cross-linked hydrogels. By "biocompatible" ismeant that the materials will meet the criteria in standard # ISO10993-1 (International Organization for Standardization, Geneva,Switzerland). By "resorbable," it is meant that the compositions willdegrade or solubilize, when placed directly into a target site in apatient's body (and not protected within an implant device such as abreast implant), over a time period of less than one year, usually from1 day to 120 days. A particular protocol for measuring resorption anddegradation is set forth in the Experimental section hereinafter. By"molecular cross-linked", it is meant that the materials comprisepolymer molecules (i.e. individual chains) which are attached by bridgescomposed of either an element, a group, or a compound, where thebackbone atoms of the polymer molecules are joined by chemical bonds.Cross-linking may be effected in a variety of ways, as will be describedin greater detail below.

By "hydrogel," it is meant that the composition comprises a single phaseaqueous colloid in which a biologic or non-biologic polymer, as definedin more detail below, absorbs water or an aqueous buffer. The hydrogelcomprises multiple sub-networks, where each sub-network is a molecularcross-linked hydrogel having dimensions which depend on the degree ofhydration and are within the ranges set forth above. Preferably, thehydrogels will have little or no free water, i.e. water cannot beremoved from the hydrogel by simple filtration.

By "percent swell," it is meant that the dry weight is subtracted fromthe wet weight, divided by the dry weight and multiplied by 100, wherewet weight is measured after the wetting agent has been removed ascompletely as possible from the exterior of the material, e.g. byfiltration, and where dry weight is measured after exposure to anelevated temperature for a time sufficient to evaporate the wettingagent, e.g., 2 hours at 120° C.

"Equilibrium swell" is defined as the percent swell at equilibrium afterthe polymeric material has been immersed in a wetting agent for a timeperiod sufficient for water content to become constant, typically 18 to24 hours.

"Target site" is the location to which the hydrogel material is to bedelivered. Usually, the target site will be the tissue location ofinterest, but in some cases the hydrogel may be administered ordispensed to a location near the location of interest, e.g. when thematerial swells in situ to cover the location of interest.

The hydrogels of the present invention may be formed from biologic andnon-biologic polymers. Suitable biologic polymers include proteins, suchas gelatin, soluble collagen, albumin, hemoglobin, casein, fibrinogen,fibrin, fibronectin, elastin, keratin, laminin, and derivatives andcombinations thereof. Particularly preferred is the use of gelatin orsoluble non-fibrillar collagen, more preferably gelatin, and exemplarygelatin formulations are set forth below. Other suitable biologicpolymers include polysaccharides, such as glycosaminoglycans (e.g.hyaluronic acid and chondroitin sulfate), starch derivatives, xylan,cellulose derivatives, hemicellulose derivatives, agarose, alginate,chitosan, and derivatives and combinations thereof. Suitablenon-biologic polymers will be selected to be degradable by either of twomechanisms, i.e. (1) break down of the polymeric backbone or (2)degradation of side chains which result in aqueous solubility. Exemplarynonbiologic hydrogel-forming polymers include synthetics, such aspolyacrylates, polymethacrylates, polyacrylamides, polyvinyl resins,polylactide-glycolides, polycaprolactones, polyoxyethylenes, andderivatives and combinations thereof.

The polymer molecules may be cross-linked in any manner suitable to forman aqueous hydrogel according to the present invention. For example,polymeric molecules may be cross-linked using bi- or poly-functionalcross-linking agents which covalently attach to two or more polymermolecules chains. Exemplary bifunctional cross-linking agents includealdehydes, epoxies, succinimides, carbodiimides, maleimides, azides,carbonates, isocyanates, divinyl sulfone, alcohols, amines, imidates,anhydrides, halides, silanes, diazoacetate, aziridines, and the like.Alternatively, cross-linking may be achieved by using oxidizers andother agents, such as periodates, which activate side-chains or moietieson the polymer so that they may react with other side-chains or moietiesto form the cross-linking bonds. An additional method of cross-linkingcomprises exposing the polymers to radiation, such as gamma radiation,to activate the side polymer to permit cross-linking reactions.Dehydrothermal cross-linking methods are also suitable. Dehydrothermalcross-linking of gelatin can be achieved by holding it at an elevatedtemperature, typically 120° C., for a period of at least 8 hours.Increasing the extent of cross-linking, as manifested in a decline inpercent swell at equilibrium, can be achieved by elevating the holdingtemperature, extending the duration of the holding time, or acombination of both. Operating under reduced pressure can accelerate thecross-linking reaction. Preferred methods for cross-linking gelatinmolecules are described below.

Optionally, the molecular cross-linked hydrogel may include aplasticizer to increase the malleability, flexibility, and rate ofdegradation of the hydrogel. The plasticizer may be an alcohol, such aspolyethylene glycol, sorbitol, or glycerol, preferably beingpolyethylene glycol having a molecular weight ranging from about 200 to1000 D, preferably being about 400 D. The plasticizers will be presentin the compositions at from about 0.1% of the solids by weight to about30% of the solids by weight, preferably from 1% of the solids by weightto 5% of the solids by weight of the composition. The plasticizers areparticularly beneficial for use with hydrogels having a high solidscontent, typically above 10% by weight of the composition (withoutplasticizer).

Exemplary methods for producing molecular cross-linked gelatins are asfollows. Gelatin is obtained and placed in an aqueous buffer to form anon-cross-linked hydrogel, typically having a solids content from 1% to70% by weight, usually from 3% to 10% by weight. The gelatin iscross-linked, typically by exposure to either glutaraldehyde (e.g. 0.01%to 0.05% w/w, overnight at 0° to 15° C. in aqueous buffer), sodiumperiodate (e.g. 0.05 M, held at 0° C. to 15° C. for 48 hours) or1-ethyl-3-(3-dimethylaminopropyl) carbodiimide ("EDC") (e.g., 0.5% to1.5% w/w, overnight at room temperature), or by exposure to about 0.3 to3 megarads of gamma or electron beam radiation. Alternatively, gelatinparticles can be suspended in an alcohol, preferably methyl alcohol orethyl alcohol, at a solids content of 1% to 70% by weight, usually 3% to10% by weight, and cross-linked by exposure to a cross-linking agent,typically glutaraldehyde (e.g., 0.01% to 0.1% w/w, overnight at roomtemperature). In the case of aldehydes, the pH should be held from about6 to 11, preferably from 7 to 10. When cross-linking withglutaraldehyde, the cross-links are formed via Schiff bases which may bestabilized by subsequent reduction, e.g. by treatment with sodiumborohydride. After cross-linking, the resulting granules may be washedin water and optionally rinsed in an alcohol, dried and resuspended to adesired degree of hydration in an aqueous medium having a desired bufferand pH. The resulting hydrogels may then be loaded into the applicatorsof the present invention, as described in more detail hereinafter.Alternatively, the hydrogels may be mechanically disrupted prior to orafter cross-linking, also as described in more detail hereinafter.

Exemplary methods for producing molecular cross-linked gelatincompositions having equilibrium percent swells in the range from about400% to about 1300%, preferably 600% to 950%, are as follows. Gelatin isobtained and placed in an aqueous buffer (typically at a pH of 6 to 11,preferably at a pH between 7 and 10) containing a cross-linking agent insolution (typically glutaraldehyde, preferably at a concentration of0.01% to 0.1% w/w) to form a hydrogel, typically having a solids contentfrom 1% to 70% by weight, usually from 3% to 10% by weight. The hydrogelis well mixed and held overnight at 0°-15° C. as cross-linking takesplace. It is then rinsed three times with deionized water, twice with analcohol (preferably methyl alcohol, ethyl alcohol, or isopropyl alcohol)and allowed to dry at room temperature. Optionally, the hydrogel may betreated with sodium borohydride to further stabilize the cross-linking.

The compositions of the present invention may be further combined withother materials and components, such as bioactive component(s) to bedelivered to the patient, viscosity modifiers, such as carbohydrates andalcohols, and other materials intended for other purposes, such as tocontrol the rate of resorption. Exemplary bioactive components include,but are not limited to, proteins, carbohydrates, nucleic acids, andinorganic and organic biologically active molecules such as enzymes,antibiotics, antineoplastic agents, bacteriostatic agents, bacteriocidalagents, antiviral agents, hemostatic agents, local anesthetics,anti-inflammatory agents, hormones, antiangiogenic agents, antibodies,neurotransmitters, psychoactive drugs, drugs affecting reproductiveorgans and oligonucleotides, such as antisense oligonucleotides. Suchbioactive components will typically be present at relatively lowconcentrations, typically below 10% by weight of the compositions,usually below 5% by weight, and often below 1% by weight. Two or more ofsuch active agents may be combined in a single composition and/or two ormore compositions may be used to deliver different active componentswhere said components may interact at the delivery site.

Exemplary hemostatic agents include thrombin, fibrinogen and clottingfactors. Hemostatic agents like thrombin may be added in concentrationsranging from 50 to 10,000 Units thrombin per ml hydrogel, preferablyfrom about 100 Units thrombin per ml hydrogel to about 1000 Unitsthrombin per ml hydrogel.

The molecular cross-linked hydrogels of the present invention can bemechanically disrupted at the time they are delivered to a target siteby extrusion through an orifice or other flow restriction, or they canbe mechanically disrupted in a batch process prior to delivery to atarget site. The primary purpose of this mechanical disruption step isto create multiple subunits of hydrogel having a size which enhances theability to fill and pack the space to which it is being delivered.Another purpose of the mechanical disruption is to facilitate passage ofthe hydrogel down small diameter tubes, cannulas, and/or otherapplicators to the target site. Without mechanical disruption, themolecular cross-linked hydrogels will have difficulty conforming to andfilling the irregularly target spaces which are being treated, e.g.intravertebral spaces in the spinal column, tissue divots, percutaneoustissue tracks, and the like. By breaking the hydrogel down to smallersized sub-units, such spaces can be filled much more efficiently whileretaining the mechanical integrity and persistence of the cross-linkedhydrogel which are essential for it to act as an anti-adhesive agent,tissue filler, or the like. Surprisingly, it has been found that asingle manual extrusion of the composition, typically using a syringehaving an orifice in size in the range from 0.01 mm to 5.0 mm,preferably from 0.1 mm to 2.5 mm, provides the proper amount ofmechanical disruption to enhance the hydrogel properties as describedabove.

Alternatively, the hydrogel compositions of the present invention may bemechanically disrupted prior to their final use or delivery. Molecularcross-linking of the polymer chains of the hydrogel can be performedbefore or after its mechanical disruption. The hydrogels may bemechanically disrupted in batch operations, such as mixing, so long asthe hydrogel composition is broken down into sub-units having a size inthe 0.01 mm to 5.0 mm range set forth above. When the hydrogelcomposition is disrupted prior to use, the hydrogel can be applied oradministered by techniques other than extrusion e.g. using a spatula,spoon, or the like. Other batch mechanical disruption processes includepumping through a homogenizer or mixer or through a pump whichcompresses, stretches, or shears the hydrogel to a level which exceeds afractural yield stress of the hydrogel. In some cases, extrusion of thepolymeric composition causes the hydrogel to be converted from asubstantially continuous network, i.e. a network which spans thedimensions of the original hydrogel mass, to a collection ofsub-networks or sub-units having dimensions in the ranges set forthabove. In other cases it may be desirable to partially disrupt thehydrogel compositions prior to packaging in the syringe or otherapplicator. In such cases, the hydrogel material will achieve thedesired sub-unit size prior to final extrusion.

In a presently preferred embodiment, the polymer may be initiallyprepared (e.g. by spray drying) and/or be mechanically disrupted priorto being cross-linked, often usually prior to hydration to form ahydrogel. The polymer may be provided as a finely divided or powdereddry solid which may be disrupted by further comminution to provideparticles having a desired size, usually being narrowly confined withina small range. Further size selection and modification steps, such assieving, cyclone classification, etc., may also be performed. For theexemplary gelatin materials described hereinafter, the dry particle sizeis preferably in the range from 0.01 mm to 1.5 mm, more preferably from0.05 mm to 1.0 mm. An exemplary particle size distribution will be suchthat greater than 95% by weight of the particles are in the range from0.05 mm to 0.7 mm. Methods for comminuting the polymeric startingmaterial include homogenization, grinding, coacervation, milling, jetmilling, and the like. Powdered polymeric starting materials may also beformed by spray drying. The particle size distribution may be furthercontrolled and refined by conventional techniques such as sieving,aggregation, further grinding, and the like.

The dry powdered solid may then be suspended in an aqueous buffer, asdescribed elsewhere herein, and cross-linked. In other cases, thepolymer may be suspended in an aqueous buffer, cross-linked, and thendried. The cross-linked, dried polymer may then be disrupted, and thedisrupted material subsequently resuspend in an aqueous buffer. In allthe cases, the resulting material comprises a cross-linked hydrogelhaving discrete sub-networks having the dimensions set forth above.

The compositions of the present invention, after mechanical disruption,will be resorbable, i.e., they will biodegrade in the patient's body, ina period of less than one year, usually from 1 to 120 days, preferablyfrom 1 to 90 days, and more preferably from 2 to 30 days following theirinitial application. This is particularly true when the materials areused for preventing post-surgical and other adhesions, where a barrieris necessary between the healing tissue surfaces only for so long as thetissue is healing. Techniques for measuring the length of time requiredfor resorption are set forth in Example 11 in the Experimental sectionbelow. In other cases, such as when the compositions are containedwithin an implantable device, such as a breast implant, resorption ofthe material will be prevented by the membrane or other mechanicalbarrier surrounding the compositions (unless the integrity of thebarrier is broken).

Referring now to FIG. 1, a method for preventing adhesions following alaminectomy procedure will be described. A syringe 10 containing theresorbable molecular cross-linked hydrogel G of the present invention isused to apply the hydrogel in such a manner that all exposed dura in avertebrae VB is covered. Usually, the hydrogel will be resorbed over atime period in the range from 7 to 60 days.

Referring now to FIGS. 2A and 2B, the molecular cross-linked hydrogelsof the present invention may also be used to fill divots D in softtissue T. A syringe 50 comprising a barrel 52, plunger 54 and cannula 56contains the molecular cross-linked hydrogel in the interior of thebarrel 52. The hydrogel G is extruded through the cannula 56 bydepressing the plunger 54 in a conventional manner. Sufficient hydrogelis extruded to fill the divot, as shown in FIG. 2B. Preferably, apartially hydrated hydrogel which will swell further upon exposure tothe moist tissue environment will be used. It may be desirable to placea patch P over the exposed surface of the hydrogel, as shown in FIG. 2B.The patch may be an adhesive or other conventional self-securing patch.Preferably, however, the patch comprises a collagen, gelatin, or otherfilm that may be immobilized by applying energy e.g. optical or radiofrequency energy as described in published PCT applications WO 96/07355and WO 92/14513.

Referring now to FIGS. 3A and 3B, compositions and methods of thepresent invention may also be used to fill percutaneous tissue tracts TTwhich were formed through overlying tissue to access blood vessels BV. Abarrier element 70 may be placed along the inner wall of the bloodvessel at the distal end of the tissue tract TT. Filament 72 may be usedto hold the barrier element 70 in place. A syringe 74 comprising abarrel 76, plunger 78, and cannula 80 is then used to extrude themolecular cross-linked hydrogel material of the present invention intothe tissue tract over the barrier element 70. The hydrogel G will beused to fill the entire interior volume of the tissue tract TT, as shownin FIG. 3B, and will preferably be partially hydrated to permitpost-placement swelling as described above. Optionally, a patch or othercover may be placed over the exposed surface of the tissue tract (notshown). The barrier element 70 may then be removed.

Referring now to FIG. 4, the present invention comprises kits includingthe hydrated, partially hydrated, and/or non-hydrated polymericcompositions described above packaged in a suitable container, usuallywith written instructions for use. For example, the composition may bepackaged in an applicator 90 which contains the pre-extruded molecularcross-linked hydrogel of the present invention. The applicator may takea wide variety of forms, including syringes as previously described. InFIG. 4, the applicator 90 comprises a tube 92 having a neck 94 whichdefines an extrusion orifice. The hydrogel is contained within the tubeand may be extruded through the neck 94 by squeezing the tube. Theapplicator 90 is preferably contained in a sterile package 96. Thesterile package may take a variety of forms, and is illustrated as anenvelope comprising a backing sheet and a clear plastic cover. Suchpackages may be sterilized in a conventional manner. Optionally, theradiation used to cross-link the hydrogel may also be used to sterilizethe entire package. The instructions for use may be printed on thepackaging or provided on a separate sheet placed in the package.

The present invention may also be used to inhibit bleeding (causehemostasis) on an abraded or damaged tissue surface, e.g., any organsurface including the liver, spleen, heart, kidney, intestine, bloodvessels, vascular organs, and the like. A syringe containing theresorbable molecular cross-linked hydrogel combined with a hemostasisagent is used to apply the hydrogel to the abraded or damaged tissuesite. The hydrogel is applied so that the actively bleeding abraded ordamaged area is completely covered with the resorbable molecularcross-linked hydrogel. Suitable hemostatic agents include thrombin,fibrinogen, and other clotting factors, as described for example in U.S.Pat. Nos. 5,411,885; 4,627,879; 4,265,233; 4,298,598; 4,362,567;4,377,572; and 4,442,655, the disclosures of which are incorporatedherein by reference. Conveniently, catalytic components of thehemostasis agent, e.g. thrombin, may be combined in the syringeimmediately prior to use so that their combined activities are preserveduntil applied to the tissue.

When used in regions surrounding nerves and other sensitive bodystructures, it is preferable to employ fully hydrated hydrogels (i.e.with >95% of hydration at equilibrium swell) in order to avoid damage tothe nerves from swelling in an enclosed environment.

The following examples are offered by way of illustration, not by way oflimitation.

EXPERIMENTAL Example 1: Materials and Methods for Production of aFragmented Polymeric Product

Fragmented polymeric compositions are generally prepared as follows:

Using pyrogen-free glassware and distilled water throughout, food gradegelatin (300 Bloom, Atlantic Gelatin, General Foods Corp., Woburn,Mass.) at 10% solids was allowed to swell in 0.1 N aq. sodium hydroxideand 0.05 sodium periodate and held at 0° C. to 8° C. for 2-3 days. Theswollen granules were washed in distilled water until the pH reached 8.The neutralized swollen granules were dried in a laminar flow hood andre-suspended in 0.05 M sodium phosphate, 0.15 M sodium chloride, pH7.2+/-0.2, at 10% solids. The composition was then loaded into 3.0 ccsyringes and irradiated at 3.0 megarad with electron beam to sterilize.

Example 2: Materials and Methods for Production of a FragmentedPolymeric Product

Gelatin (300 Bloom, Atlantic Gelatin, General Foods Corp., Woburn,Mass.) was allowed to swell in an aqueous buffer (e.g. 0.05 M sodiumphosphate, 0.15 M sodium chloride, pH 7.2+/-0.2) at 1-10% solids and wascross-linked by either glutaraldehyde (0.01-0.05%, w/w, overnight, roomtemperature), by sodium periodate (0.05 M, 0° C. to 8° C., 48 hours) orby 0.3-3.0 megarads of gamma or electron beam irradiation. The hydrogelswere then extruded from a syringe using normal manual pressure.

Example 3: Materials and Methods for Production of a FragmentedPolymeric Product

Gelatin (300 Bloom, Atlantic Gelatin, General Foods Corp., Woburn,Mass.) was allowed to swell in distilled water at 1-10% solids (w/w)chilled to 5° C. The resultant hydrogel was fragmented by stirring withan impeller driven by a motor. Then, sodium periodate and sodiumhydroxide were added and mixed to achieve 0.05 M sodium periodate and0.10 M sodium hydroxide. The chilled mixture was held at 0° C. to 8° C.for 2-3 days. The cross-linked hydrogel fragments were then washed with5° C. water to achieve pH 8. Finally the hydrogel fragments were washedwith an aqueous buffer (e.g. 0.05 sodium phosphate and 0.15 sodiumchloride, pH 7.2+/-0.2) and left at 0° C. to 8° C. to equilibrate withthe buffer. Free buffer was decanted from the fragmented hydrogel massand the hydrogel particles were loaded into syringes and irradiated at3.0 megarads by electron beam or gamma irradiation to sterilize. Suchsterilized fragmented were then extruded directly from the syringecausing further fragmentation.

Example 4: Biocompatibility in Rabbit Model

The test material was prepared by mixing 0.5 mL of sterile saline forinjection with 5 mL of the fragmented gelatin composition as follows:The saline solution was injected into the fragmented gelatin compositioncontained in a 5 cc syringe through a dispersion needle embedded in thefragmented gelatin composition. One mL aliquots of the fragmentedgelatin composition were transferred into 1 cc syringe and a 14 gaugeneedle was attached. The entire assembly was weighed. Followingadministration of the test article, the syringe and needle assembly werere-weighed to determine the mass of the dosed compound.

A total of 14 rabbits were included in this study. All procedures wereperformed aseptically. Rabbits were clipped free of fur over theparavertebral muscles. The test material was delivered from a 1 mLsyringe with a 14 gauge needle. The needle attached to the syringecontaining the test article was inserted into the muscle at a 45° angle.An approximate 0.2 mL portion of the test material was injected into themuscle and needle withdrawn. A total of four test sites were implantedin the right paravertebral muscle of each rabbit. Additionally, a USPnegative control strip was implanted as a marker approximately 2-3 mmaway from each test site, distal to the vertebral column as compared tothe test article. In the opposite (left) muscle, four USP negativecontrol sections were implanted similarly as performed for the markers.

Observations included daily health checks, adverse reactions related toimplantation, morbidity and mortality. Body weights were recorded priorto implantation, at monthly intervals, and at termination. At each ofthe harvest times post-implantation; 2, 4, 6 and 13 weeks, 3 rabbitswere euthanized. A gross examination for irritation at each implant sitewas performed. The paravertebral muscles were dissected free and fixedin 10% neutral buffered formalin. Following appropriate embedding,sectioning and staining, the muscles were evaluated microscopically forevidence of irritation, presence or absence of test article and relativedegree of resorption of the implanted test article.

All animals appeared clinically normal throughout the study and gainedappropriate weights during the course of the study. At 2 weeks, theinflammatory reaction around the injected test material was localized,with very little extension of this inflammation observed beyond the testmaterial. At four weeks, the inflammatory reaction around the injectedtest material was localized, with very little extension of thisinflammation observed beyond the test material. At four weeks, there wasa minimal to mild inflammatory and fibrotic reaction observed at thetest sites, which resolved to a minimal reaction at six weeks. Bythirteen weeks the inflammatory response was characterized as extremelyminimal. The test material was considered to be a non-irritant, comparedto the USP negative control material at six and thirteen weeks postimplantation.

Example 5: Vessel Plug

This study demonstrated the effectiveness of the fragmented polymericcomposition to seal a vessel puncture. The femoral artery of a farmgrade Hampshire/Yorkshire cross pig (Pork Power Farms, Turlock, Calif.)was identified and cannulated using a needle (SmartNeedle™,CardioVascular Dynamics, Irvine, Calif.). After the guide wire wasplaced, a 9 French dilator was used to create a tunnel to the vessel andenlarge the femoral artery opening. The dilator was removed and a 7French sheath was introduced into the femoral artery. The guide wire wasthen removed. Positioning was checked by withdrawing blood into thesheath side arm. Pulsatile arterial bleeding was also observed at thepoint of insertion of sheath at the skin incision. As the sheath wasremoved, a 18 gauge Teflon catheter tip attached to a hypodermic syringewas used to introduce the fragmented gelatin composition of Example 1into the tunnel. No bleeding was observed at the point of exitdemonstrating the effectiveness of the fragmented gelatin composition insealing the vessel puncture site and surrounding tissue.

Example 6: Fragmented Polymeric Composition as a Carrier

This study demonstrated the effectiveness of the fragmented polymericcomposition of Example 1 as a carrier to fill and seal a tissue divot inthe liver. Three wounds (2 tissue divots and 1 tissue puncture) wereinduced in the liver of a farm grade Hampshire/Yorkshire cross pig (PorkPower Farms, Turlock, Calif.).

Liver tissue divot #1 was actively bleeding following the surgicalcreation of a tissue divot. A syringe, containing approximately 1 ml offragmented gelatin composition containing approximately 500 U ofthrombin (500 to 1000 units/ml) was extruded from a syringe and appliedto completely fill the tissue defect. After 2-3 minutes, a blood clotformed causing immediate cessation of bleeding. When the appliedcomposition was grasped with forceps, it appeared to adhere quite wellto the tissue and had good integrity. The sealant was manuallychallenged and no additional bleeding was observed.

Liver tissue divot #2 was actively bleeding following the surgicalcreation of a tissue divot. Approximately 1 ml of fragmented gelatincomposition containing thrombin (approximately 500 units/ml) wasextruded from a syringe and applied to completely fill the tissuedefect. A Rapiseal™ patch (Fusion Medical Technologies, Inc., MountainView, Calif.) was applied over the filled defect using an argon beamcoagulator (Valleylab, Boulder, Colo., or Birtcher Medical Systems,Irvine, Calif.,). Immediate cessation of bleeding occurred.

Liver puncture #1, was actively bleeding following the surgical creationof a blunt puncture. Approximately 0.8 ml of fragmented gelatincomposition containing thrombin (approximately 500 units/ml) wasextruded from a syringe and applied to completely fill the tissuedefect. Approximately 2 minutes following the delivery of the fragmentedgelatin composition, all bleeding stopped.

Spleen puncture #1 was actively bleeding following the surgical creationof a blunt puncture. Approximately 0.8 ml of fragmented gelatincomposition containing thrombin (approximately 500 units/ml) wasextruded from a syringe and applied to completely fill the tissuedefect. Approximately 2 minutes following the delivery of the fragmentedgelatin composition, all bleeding stopped.

In the above four examples, the delivery system used was a 3 cc syringe(Becton Dickinson, Franklin Lakes, N.J.). It contained the fragmentedgelatin composition of example 1.

A material according to the present invention for filling tissue divotsand other defects could be prepared as follows. A thrombin solution (0.5ml; 4,000 to 10,000 U/ml) is added to 4.5 ml of flowable hydrogel toproduce 5 ml of hydrogel containing 400 to 1000 U/ml thrombin. Thehydrogel can be used in any convenient amount, e.g. 0.5 ml to 5 ml.

Example 7: Fragmented Polymeric Composition as a Tissue Filler andAnastomic Sealant

This study demonstrated the effectiveness of the fragmented gelatincomposition as a wound closure system that fills and seals tissuedefects. Four tissue divots were surgically induced, 1 in the lung, 2 inthe liver and 1 in the spleen of a farm grade Hampshire/Yorkshire crosspig (Pork Power Farms, Turlock, Calif.).

On the lung, following the surgical creation of the tissue divot, an airleak was observed. Approximately 1 ml of the fragmented gelatincomposition of Example 1 was extruded as from a syringe and applied tocompletely fill the tissue defect. A Rapiseal™ patch (Fusion MedicalTechnologies, Inc., Mountain View, Calif.) was applied using an argonbeam coagulator (Valleylab, Boulder, Colo., or Birtcher Medical Systems,Irvine, Calif.,). Immediate cessation of the air leak occurred. When theapplied patch was grasped with forceps, it appeared to adhere quite wellto the tissue and had good integrity. The fragmented gelatin compositionwas challenged by ventilating the lung to a pressure of 28 cm water. Noair leak was observed.

On the liver, following the surgical creation of the tissue divot,excessive bleeding was observed. Approximately 1 ml of fragmentedgelatin composition was extruded from a syringe and applied tocompletely fill the tissue defect. The fragmented composition swelledand adequately stopped the bleeding although some seepage bleeding wasobserved.

On the liver, following the surgical creation of the tissue divot,excessive bleeding was observed. Approximately 1 ml of fragmentedgelatin composition was extruded from a syringe and applied tocompletely fill the tissue defect. A Rapiseal™ patch (Fusion MedicalTechnologies, Inc., Mountain View, Calif.) was applied using an argonbeam coagulator (Valleylab, Boulder, Colo., or Birtcher Medical Systems,Irvine, Calif.,). Immediate cessation of the bleeding occurred. When theapplied patch was grasped with forceps, it appeared to adhere quite wellto the tissue and had good integrity.

Spleen puncture #1 was actively bleeding following the surgical creationof a blunt puncture. Approximately 0.8 ml of fragmented gelatincomposition was extruded from a syringe and applied to completely fillthe tissue defect. Approximately 2 minutes following the delivery of thefragmented gelatin composition, all bleeding stopped.

A female juvenile farm grade goat (Clovertop Dairy, Madera, Calif.) wasused under appropriate anesthesia. The right cartoid artery was exposed.The vessel was carefully dissected to remove any connective tissue. Thevessel was clamped using atraumatic vascular clamps, separated by adistance of approximately 2-3 cm. The vessel was dissected using astandard scalpel blade to expose 2 free vessels ends. An end-to-endanastomosis was created using 6-0 prolene suture in an interruptedfashion. Following completion of the anastomoses, the clamps werereleased. Bleeding was observed at the anastomotic site. Approximately 2cc of the fragmented gelatin composition containing thrombin(approximately 500 units/ml) was extruded from a syringe around theanastomoses. Gauze was placed against the composition. Approximately 3minutes after the application of the fragmented gelatin composition, allbleeding was observed to have ceased. The incision was appropriatelyclosed and the animal was allowed to recover for subsequent follow-up.

Example 8: Materials and Methods for Determining Force Necessary toRelease Aqueous Phase

Two disks were cut from a filter mesh of sufficient pore size to retainthe sample under test. The disks were of approximately the same diameteras the inside of the barrel of a 5 ml syringe (Becton Dickinson,Franklin Lakes, N.J.). The plunger was removed from the 5 ml syringe andthe two mesh disks were inserted and pushed into place with the plunger.The plunger was replaced and the syringe placed into an assemblyallowing the syringe plunger to be depressed by the force gauge of aChatillon TCD 200 Test Stand (Chatillon, Greensboro, N.C.). The forcerequired to cause the release of aqueous phase from the test materialwas then determined.

A 51 μm stainless steel mesh was used to retain the test materials. Theapplication of 50 lbs. force to 2 ml of the material of Example 4(above) mixed with reconstituted thrombin (Thrombin-JMI™, GenTrac, Inc.,Middelton, Wis.) according to the package insert was insufficient tocause the release of any free liquid, nor any noticeable collapse of thematerial.

A sterile absorbable gelatin sponge (2.5 ml; Gelfoam®, the UpJohn Co.,Kalamazoo, Mich.) was soaked in reconstituted thrombin (Thrombin-JMI™,GenTrac, Inc., Middelton, Wis.) according to the package insert andinserted into the same apparatus as above. The application of less than1 lb. of pressure caused the release of almost all of the aqueous phaseand the collapse of the Gelfoam material to approximately 0.5 mL.

Example 9: Materials and Methods of Ascorbate Addition to Hydrogel Priorto Irradiation

Gelatin particles (300 Bloom, Atlantic Gelatin, General Foods Corp.,Woburn, Mass.) were suspended at 5%-15% by weight in methyl alcohol(Aldrich, Milwaukee, Wis.) containing 0.01%-0.1% by weightglutaraldehyde (Sigma, St. Louis, Mo.) and stirred overnight at ambienttemperature. Alternatively, gelatin particles, obtained from an extractof calf hide (Spears Co., PA) were suspended at 5%-15% by weight inaqueous buffer at pH 9 containing 0.01%-0.l% by weight glutaraldehyde(Sigma) to form a hydrogel that was well-mixed and refrigeratedovernight. The cross-linked gelatin fragments were then rinsed threetimes with alcohol and dried at ambient temperature. Equilibriumswelling for the rinsed, cross-linked gelatin was then measured, and 0.5g-1.0 g portions of this material were packed into 5 cc syringes. 3.0ml-4.5 ml of aqueous buffer containing ascorbic acid or a salt ofascorbic acid, e.g. 0.02 M sodium phosphate (J.T. Baker, Phillipsburg,N.J.), 0.15 M sodium chloride (VWR, West Chester, Pa.), 0.005 M sodiumascorbate (Aldrich), pH 7.0, was added to the syringes containingcross-linked gelatin using a second syringe and a three-way stopcock,with care taken not to introduce extraneous air into the syringes, toform a hydrogel within several syringes. Alternatively, an aqueousbuffer that did not contain ascorbic acid or a salt of ascorbic acid butwas otherwise of similar composition and pH was added to other syringescontaining cross-linked gelatin to form a hydrogel within them. Thehydrogel-containing syringes were then gamma-irradiated underrefrigerated conditions at 3.0±0.3 megarads. Equilibrium swell wasmeasured for the hydrogel contained within the syringes afterirradiation. Hydrogels that were formed using buffers that containedascorbic acid or a salt of ascorbic acid generally maintained values forequilibrium swell upon irradiation within ±20%, and usually ±10%, of thevalue prior to irradiation, while hydrogels that were formed usingbuffers not containing ascorbic acid or a salt of ascorbic acidexperienced a decrease in equilibrium swell of 25-30% of its value priorto irradiation.

Example 10: Materials and Methods of Cross-Linking and Measuring PercentSwell

Gelatin particles were allowed to swell in an aqueous buffer (e.g., 0.2M sodium phosphate, pH 9.2) containing a cross-linking agent (e.g.,0.005-0.5% by weight glutaraldehyde). The reaction mixture was heldrefrigerated overnight and then rinsed three times with deionized water,twice with ethyl alcohol, and allowed to dry at ambient temperature. Thedried, cross-linked gelatin was resuspended in an aqueous buffer at alow solids concentration (2-3%) at ambient temperature for a fixedperiod of time. Buffer was in substantial excess of the concentrationneeded for equilibrium swelling, and two phases (a hydrogel phase and abuffer) were present. The suspension containing wet hydrogel was thenfiltered by applying vacuum on a 0.8 μm nominal cut-off filter membrane(Millipore, Bedford, Mass.). After removal of extraneous buffer, thecombined weight of the retained wet hydrogel and wet filter membrane wasrecorded. The hydrogel and membrane were then dried at approximately120° C. for at least two hours, and the combined weight of the driedhydrogel residue and dried filter membrane was recorded. Severalmeasurements of samples of wet filter membrane without hydrogel residueand dried filter membrane without hydrogel were also performed and wereused to deduce the net weight of wet hydrogel and dry hydrogel. "Percentswell" was then calculated as follows: ##EQU1##

Swell measurements were conducted at least in triplicate and averagedfor a given sample of gelatin. The value of percent swell for samplesresuspended in buffer for 18-24 hr prior to measuring wet weight wasdefined as "equilibrium swell."

The resulting cross-linked gelatin materials displayed equilibrium swellvalues in the range from 400% to 1300%. The degree of equilibrium swelldepended on the method and extent of cross-linking.

Example 11: Degradation

Thirty rabbits (15 untreated control animals and 15 animals treated withfragmented gelatin composition) underwent surgery to mimic splenicinjury and bleeding. A lesion on the spleen was created by making acontrolled wound with a 6 mm biopsy punch. In the "Treated" group, theexperimentally created injury was immediately treated with thefragmented gelatin composition to cause hemostasis of the wound."Control" group animals were not treated during the first 7.5 minutes todemonstrate the amount of bleeding resulting from the lesion. At 7.5minutes from the time the injury was caused, the fragmented gelatincomposition was then used to stop bleeding from the lesion to preventspontaneous exsanguination and death of the animal. All animals wereallowed to recover. Ten animals each were euthanized on Days 14 and 28post-surgery. The final necropsy date for the remaining animals wasdetermined after the Day 28 animals were evaluated. In animals harvestedat the Day 28 time point it was difficult to determine via grossexamination if the test material was present or not, therefore half ofthe remaining animals were harvested at Day 42 and the other half at Day56. At the time of necropsy, the site of the splenic lesion and theperitoneal cavity were evaluated macroscopically. Presence of fragmentedgelatin composition in the peritoneal cavity away from the site ofplacement was noted and evaluated, as well as its presence or absence atthe splenic lesion. The presence or absence of postoperative adhesionsat the site of the splenic lesion was also evaluated and noted. Thespleen was carefully dissected and processed for histological evaluationof biocompatibility and biodegradation.

The application of the fragmented gelatin composition to the surgicallycreated wounds on the spleen resulted in good hemostatic tamponade.Following application of the fragmented gelatin composition at the timeof surgery, rabbits were survived for 14, 28, 42, and 56 dayspostoperatively. One rabbit died of unrelated pneumonia at Day 5postoperatively and the spleen was not harvested for histopathologicalexamination.

At necropsy, the site of the splenic lesion as well as the peritonealcavity in general were evaluated grossly. Presence of the fragmentedgelatin composition in the peritoneal cavity away from the site ofplacement was evaluated, as well as the presence or absence of thefragmented gelatin composition at the splenic lesion. The presence orabsence of adhesions at the site of the splenic lesion were evaluatedand noted. The spleen was carefully dissected and processed forhistological evaluation.

Grossly, the site of the splenic lesion was visible in all animals, atall time points. Macroscopically, the fragmented gelatin composition wasabsent in two of the ten Day 14 animals. At all other time points it wasnot possible to identify the fragmented gelatin compositionmacroscopically. The macroscopic absence of the hydrogel material asmeasured in this rabbit model defines the degradation of the hydrogel asthat term is used herein and in the claims.

In three of ten animals sacrificed at 14 days postoperatively, smallamounts of the fragmented gelatin composition were found free-floatingin the abdominal cavity. This most likely represents the excess materialthat had migrated from its placement site at the splenic lesion. In nocase where this material was found away from the splenic lesion wasthere any evidence of tissue reaction from the visceral surfaces or theomentum. No material was found away from the site of the splenic lesionin animals that were harvested at any other time point.

No postoperative adhesions associated with the fragmented gelatincomposition material were noted at the site of the splenic lesion in anyanimal. In all animals, as expected, there was omentum attached to thesite of the splenic lesion. Other adhesions involving the spleen wererare, and when noted were incidental and usually associated with theincision of the body wall.

The fragmented gelatin composition was absent macroscopically andmicroscopically in two of the ten animals from the 14 day time point. At28 days post-implant, the fragmented gelatin composition was not visibleon gross observation and microscopically was completely absent in fiveout of ten rabbits examined and present in minimal amounts in theremaining animals, showing that the fragmented gelatin composition wascomposition was essentially biodegraded by 28 days. The fragmentedgelatin composition was completely absent in all five animals examinedat 42 days post-implant and was found in minimal amounts in only one offour rabbits examined at 56 days post-implant. Healing of the splenicwound was proceeding in a normal fashion at Day 42 and more so at Day56.

Example 12: Fragmented Polymeric Product Composed of GelatinCross-Linked Using EDC

Gelatin (Atlantic Gelatin, General Foods Corp., Woburn, Mass.) wasallowed to dissolve in distilled water at 1-10% solids (w/w) (morepreferably at 8%) at 70° C. 1-Ethyl-3-(3dimethylaminopropyl)carbodiimide (EDC) (Sigma, St. Louis, Mo.) at 0.2%-3.5% (more preferably0.2%-0.3%) was then added. The resultant hydrogel formed on stirring wasleft at room temperature for one hour. The hydrogel was dried using aFreezone 12 freeze dry system, (Labconco, Mo.) and ground finely using aWaring Blender model No. 31BC91 (VWR, Willard, Ohio). The driedpolymeric composition was then loaded into syringes and equilibratedwith buffer. The equilibrium swell was determined to be at least 1000%according to the method described in Example 10. The results are shownin Table 1.

                  TABLE 1                                                         ______________________________________                                        Gelatin (mg) EDC          Swell (%)                                           ______________________________________                                        500 (8%)     13.5 mg (0.25%)                                                                            1080                                                500 (8%)     13.5 mg (0.25%)                                                                            1126                                                100 (7.4%)   0.945 mg (0.35%)                                                                           1620                                                100 (7.4%)   9.45 mg (3.5%)                                                                             1777                                                ______________________________________                                    

Example 13: Fragmented Polymeric Product Composed of Gelatin andPoly(L)Glutamic Acid, Cross-Linked Using EDC

Gelatin (Atlantic Gelatin, General Foods Corp., Woburn, Mass.) wasallowed to dissolve in distilled water at 1-10% solids (w/w) (morepreferably at 6-8%) at 70° C. 0-10% (w/w) (more preferably 2-5%)Poly(L)glutamic acid (PLGA) (Sigma, St. Louis, Mo.) and1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) (Sigma) at 0.2-3.5%(preferably 0.2-0.4%) were then added. The resultant hydrogel formed onstirring was left at room temperature for one hour. The hydrogel wasallowed to swell in excess saline for a fixed period of time (preferably20 hr.) The hydrogel was then filtered by applying vacuum on a filtermembrane (Millipore, Bedford, Mass.). The equilibrium swell wasdetermined to be at least 1500% according to the method described inExample 10. The results are shown in Table 2.

                  TABLE 2                                                         ______________________________________                                        Gelatin (mg)                                                                            PLGA (mg)   EDC           Swell (%)                                 ______________________________________                                        375  (6%)     125    (2%)   13.5 mg                                                                              (.25%) 1510                                375  (6%)     125    (2%)   13.5 mg                                                                              (.25%) 1596                                250  (4%)     250    (4%)   13.5 mg                                                                              (.25%) 2535                                250  (4%)     250    (4%)   13.5 mg                                                                              (.25%) 2591                                250  (4%)     250    (4%)   13.5 mg                                                                              (.25%) 2548                                250  (4%)     250    (4%)   13.5 mg                                                                              (.25%) 2526                                200  (3.2%)   300    (4.8%) 13.5 mg                                                                              (.25%) 2747                                200  (3.2%)   300    (4.8%) 13.5 mg                                                                              (.25%) 2677                                200  (3.2%)   300    (4.8%) 13.5 mg                                                                              (.25%) 2669                                150  (2.4%)   350    (5.6%) 13.5 mg                                                                              (.25%) 3258                                150  (2.4%)   350    (5.6%) 13.5 mg                                                                              (.25%) 3434                                150  (2.4%)   350    (5.6%) 13.5 mg                                                                              (.25%) 3275                                75   (5.5%)   25     (1.9%) 0.945 mg                                                                             (0.35%)                                                                              2437                                50   (3.7%)   50     (3.7%) 0.945 mg                                                                             (0.35%)                                                                              2616                                25   (1.9%)   75     (5.5%) 0.945 mg                                                                             (0.35%)                                                                              5383                                75   (5.5%)   25     (1.9%) 9.45 mg                                                                              (3.5%) 1976                                50   (3.7%)   50     (3.7%) 9.45 mg                                                                              (3.5%) 2925                                25   (1.9%)   75     (5.5%) 9.45 mg                                                                              (3.5%) 4798                                ______________________________________                                    

Example 14: Production of a Fragmented Polymeric Hydrogel

Bovine Corium (Spears Co. PA) was agitated in an aqueous sodiumhydroxide (Spectrum Chemical Co., CA) solution (0.1 M to 1.5 Mpreferably 0.4 to 1.2M) for a period of one to 18 hours (preferably oneto four hours) at a temperature of 2° C. to 30° C. (preferably 22° C. to30° C.). The corium slurry was then neutralized using an inorganic acidsuch as hydrochloric acid, phosphoric acid or sulfuric acid (SpectrumChemical Co., CA.) and the neutralized liquid phase was then separatedfrom the insoluble corium by filtration through a sieve. The corium wasthen washed with non-pyrogenic water and an alcohol such as isopropylalcohol (Spectrum Chemical Co., CA.). After three to twelve washes, thecorium was suspended in non-pyrogenic water and the corium, water slurrymay be then heated to 50° C. to 90° C. preferably 60° C. to 80° C. tothermally gelatinize the corium. During the gelatinization cycle, the pHof the corium, water slurry was adjusted and controlled from pH 3 to pH11, preferably pH 7 to pH 9. Also, the insoluble corium in the slurrymay be disrupted by agitation and/or homogenization. The disruption canoccur before or after the thermal gelatinization cycle. Thermalgelatinization was conducted for one to six hours. After gelatinization,the slurry was clarified by filtration. The gelatin slurry was dewateredby drying in air at 15° C. to 40° C., preferably 20° C. to 35° C. Thedry gelatin, where dry implies a moisture content less than 20% byweight, was then disrupted by grinding.

Dry gelatin was added to a cold (5° C. to 15° C.) aqueous solution ofcontaining glutaraldehyde (Amresco Inc., OH.) at 0.0025% to 0.075% byweight and at a pH between 7 and 10. The concentration of gelatin inthis solution was between 1% and 10% by weight. The glutaraldehydecross-links the gelatin granules over a period of one to 18 hours afterwhich the gelatin was separated from the aqueous phase by filtration orsedimentation. The gelatin particles were then added to an aqueoussolution containing 0.00833% to 0.0667% by weight sodium borohydride(Spectrum Chemical Co., CA.) with the gelatin concentration again beingbetween 1% and 10% by weight and the pH being between 7 and 12,preferably 7 to 9. After one to six hours, the cross-linked gelatin wasseparated from the aqueous phase by filtration or sedimentation. Thegelatin may then be resuspended in non-pyrogenic water with the gelatinconcentration being between 1% and 10% by weight to remove residualcross-linking and reducing agents followed by separation from theaqueous phase by filtration or sedimentation. Final collection of thecross-linked gelatin was done on a filter mesh or sieve and the gelatinwas given a final rinse with non-pyrogenic water. The wet, cross-linkedgelatin was then placed in a drying chamber at 15° C. to 40° C. Dry,cross-linked gelatin (i.e. cross-linked gelatin with a moisture contentbelow 20% by weight) was removed from the drying chamber and then groundusing a mechanical, grinding mill to produce a powder with a typicalparticle size distribution from 0.020 mm to 2.000 mm.

Dry, powdered, cross-linked gelatin was resuspended in a sodiumphosphate, sodium chloride and sodium ascorbate buffer at pH 5 to 8 withthe gelatin concentration being 10% to 20% by weight. The dry, powdered,cross-linked gelatin may be mixed with the buffer before dispensing thematerial into the applicator device (i.e. a syringe) or the powdered,cross-linked gelatin may be mixed with the buffer within the applicatordevice (i.e. a syringe). Additionally, a gas, such as air or nitrogen,may be dispersed with the gelatin and buffer to aid in mixing anddispensing of the material. The gas typically comprised less than 20% byvolume of the final mixture.

Powdered, cross-linked gelatin, mixed with buffer and gas within anapplicator device was then sealed in the kit to be supplied to the enduser. The kits are sterilized by irradiation with gamma-rays or anelectron beam.

Although the foregoing invention has been described in some detail byway of illustration and example, for purposes of clarity ofunderstanding, it will be obvious that certain changes and modificationsmay be practiced within the scope of the appended claims.

What is claimed is:
 1. A fragmented biocompatible hydrogel which is atleast partially hydrated and is substantially free from an aqueousphase, wherein said hydrogel comprises gelatin and will absorb waterwhen delivered to a moist tissue target site and, wherein the gel has asubunit size in the range from 0.01 mm to 5 mm when fully hydrated andan equilibrium swell from 400% to 5000%, said hydrogel being present inan applicator.
 2. The hydrogel of claim 1, having an in vivo degradationtime of less than one year.
 3. The hydrogel of any of claims 1 and 2,said hydrogel being at least partially hydrated with an aqueous mediumcomprising an active agent.
 4. The hydrogel of claim 3, wherein theactive agent is a clotting agent.
 5. The hydrogel of claim 4, whereinthe clotting agent is thrombin.
 6. A method for delivering an activeagent to a patient, said method comprising administering to a targetsite on the patient an amount of the hydrogel of claim
 3. 7. A methodfor delivering a clotting agent to a patient, said method comprisingadministering from the applicator to a bleeding target site an amount ofthe hydrogel of claim 4 sufficient to inhibit bleeding.
 8. A method fordelivering thrombin to a patient, said method comprising administeringfrom the applicator to a bleeding target site an amount of the hydrogelof claim 5 sufficient to inhibit bleeding.